Biosensor devices, systems and methods therefor

ABSTRACT

A sensing apparatus for sensing target materials including biological or chemical molecules in a fluid. One such apparatus includes a semiconductor-on-insulator (SOI) structure having an electrically-insulating layer, a fluidic channel supported by the SOI structure and configured and arranged to receive and pass a fluid including the target materials, and a semiconductor device including at least three electrically-contiguous semiconductor regions doped to exhibit a common polarity. The semiconductor regions include a sandwiched region sandwiched between two of the other semiconductor regions, and configured and arranged adjacent to the fluidic channel with a surface directed toward the fluidic channel for coupling to the target materials in the fluidic channel, and further arranged for responding to a bias voltage. The sensing apparatus also includes an amplification circuit in or on the SOI and that is arranged to facilitate sensing of the target material near the fluidic channel.

RELATED DOCUMENTS

This patent document is a continuation under 35 U.S.C. §120 of U.S.patent application Ser. No. 13/252,315 filed on Oct. 4, 2011 (U.S. Pat.No. 9,184,099) which claims benefit under 35 U.S.C. §119 to U.S.Provisional Patent Application Ser. No. 61/389,590, entitled “BiosensorDevices, Systems and Methods Therefor” and filed on Oct. 4, 2010; eachof these patent documents and the Appendices filed in the underlyingprovisional application, including the references cited therein, arefully incorporated herein by reference.

BACKGROUND

Despite technical improvements, bio-based and related sensor technologyinvolving industries, such as the semiconductor industry, has beenchallenging to implement and/or use to obtain desirable results. Forexample, sensors used in biomedicine (biosensors) can be classified intocategories including optical and electrical-based biosensors. Opticalbiosensors generally exhibit higher sensitivity and wider range ofdetection, but can suffer from lack of real time and label freedetection. Electrical biosensors, such as impedance biosensors, canaddress issues relating to optical biosensors for applications such aspoint-of-care and disease discovery, using real time, low cost, ease ofminiaturization and label-free operation. However, such electricalbiosensors have also been challenging to implement.

These and other matters have presented and continue to presentchallenges to a variety of sensor technologies, including those relatingto biomedical applications.

SUMMARY

Aspects of the present disclosure relate generally to biosensors,biosensor devices, biosensor systems and methods relating to theiroperation as discussed above.

One aspect of the present disclosure relates to a method, device and/orsystem directed to sensing a biological material using a sensor having asemiconducting channel bridge with one or more surfaces exposed forinteracting with the biological material, and which may have a localamplifier.

Certain other, more specific, aspects of the disclosure are directed toa sensing apparatus for sensing target materials including biological orchemical molecules in a fluid. One such apparatus includes a substratesuch as a semiconductor-on-insulator (SOI) structure including anelectrically-insulating layer, a fluidic channel supported by thesubstrate structure and configured to receive and pass a fluid includingthe target materials, and a semiconductor device including at leastthree electrically-contiguous semiconductor regions doped to exhibit acommon polarity. The semiconductor regions including a sandwiched regionsandwiched between two of the other semiconductor regions, andconfigured and arranged adjacent to the fluidic channel with a surfacedirected toward the fluidic channel for coupling to the target materialsin the fluidic channel, and further arranged for responding to a biasvoltage. By doping two of the other semiconductor regions at a higherconcentration than the sandwiched region, optionally with a bias voltageapplied to a reference electrode or back bias source, and in response tothe target materials in the fluidic channel, the three contiguoussemiconductor regions facilitate an interaction to sense the presence ofthe target materials in the fluidic channel by passing current in aconducting mode from one of the end electrodes to the other of the endelectrodes. Also, an amplification circuit in or on the SOI is arrangedto facilitate sensing the target material near the fluidic.

Other aspects are directed to a method, device and/or system directed tosensing biological materials using a fluidic channel to receive and passa fluid including the biological materials. A semiconductor channelbridging the microfluidic channel and having opposing surfaces exposedfor coupling to biological materials in the microfluidic channel.Biological materials may change the surface potential of thesemiconductor via a first surface of the semiconductor channel, andbiological materials may apply a bias to the semiconductor channel via asecond surface of the semiconductor channel that is opposite the firstsurface, with source and drain electrodes being connected by thesemiconductor channel. An integrated amplifier amplifies current passingbetween the electrodes via the sandwiched semiconductor regions andbiomolecules coupled or associated thereto, the current being indicativeof a conductance characteristic of the channel as coupled to biologicalmaterials.

The above overview is not intended to describe each illustratedembodiment or every implementation of the present disclosure.

DESCRIPTION OF THE DRAWINGS AND EXAMPLE EMBODIMENTS

Various example embodiments may be more completely understood inconsideration of the following detailed description in connection withthe accompanying drawings, in which:

FIG. 1A shows an example embodiment of an apparatus for sensingbiological materials including a microfluidic channel;

FIG. 1B shows another example embodiment of an apparatus for sensingbiological materials including a microfluidic channel;

FIG. 2 shows an example embodiment of a method of manufacturing anapparatus for sensing biological materials;

FIG. 3 shows an optical microscopy image of an example embodiment of anapparatus for sensing biological materials;

FIG. 4A shows an example embodiment of a biological sensor having arough surface;

FIG. 4B shows a simulated rough-surface sensor and a smooth-surfacesensor for DNA molecule concentration of a solution before and afterpassing the sensor;

FIG. 5A shows an example embodiment of electro-osmosis throughapplication of an AC signal on neighboring electrodes in an ionicsolution;

FIG. 5B shows an example embodiment of electro-osmosis andelectrophoresis application around a biological sensor;

FIG. 6A shows an example embodiment of an apparatus for sensingbiological materials including a microfluidic channel;

FIG. 6B shows an example embodiment of an apparatus for sensingbiological materials including a microfluidic channel; and

FIG. 7 is a graph showing electrical characteristics of three types ofdesigns of embodiments consistent with the present disclosure.

While the disclosure is amenable to various modifications andalternative forms, examples thereof have been shown by way of example inthe drawings and will be described in detail. It should be understood,however, that the intention is not to limit the disclosure to theparticular embodiments shown and/or described. On the contrary, theintention is to cover all modifications, equivalents, and alternativesfalling within the spirit and scope of the disclosure.

DETAILED DESCRIPTION

Various aspects of the present disclosure are directed to sensors,including biosensors implementing impedance modulation with a nanobridgetype of biosensor, for characterizing biochemical species such asantibodies and antigens, and/or for applications involving one or moreof DNA sequencing, DNA hybridization, Real time PCR, protein or otherbio-species and chemical-species detection. While the present disclosuremay not be so limited, applications thereof may be appreciated using adiscussion of example embodiments in this context.

In an example embodiment of the instant disclosure, an apparatus isconfigured and arranged to sense target materials. The target typicallyincludes biological or chemical molecules in a fluid. The device mayinclude a semiconductor-on-insulator (SOI) structure and anelectrically-insulating layer. Alternately, the device may be fabricatedon a silicon substrate without an insulating layer below the threeelectrically-contiguous semiconductor regions. In some embodiments thesilicon substrate may be of opposite polarity of the threeelectrically-contiguous semiconductor regions. In some embodiments aburied channel may be used to create a thin area a of low conductivityzone for one or more of the electrically-contiguous semiconductorregions. A fluidic channel passes and receives the fluid, including thetarget materials, and may be located on and supported by the SOIstructure. The apparatus further includes a semiconductor device whichhas at least three electrically-contiguous semiconductor regions. The atleast three electrically-contiguous semiconductor regions are doped toexhibit a common polarity. In certain more specific embodiments, thecommon polarity is of a p-type doping scheme, and in other embodiments,the common polarity is an n-type doping scheme. In other embodiment, thedifferent regions may have different doping polarities.

The semiconductor regions are characterized as having a regionsandwiched between two of the other semiconductor regions. Thesandwiched region is adjacent to the fluidic channel, and has a surfacedirected toward the fluidic channel for coupling to the target materialsthat pass through the channel. In certain specific embodiments, thesurface includes beads which may be in the fluidic channel for targetmaterial attachment. The sandwiched region is further arranged torespond to a bias voltage.

In some embodiments, the surface directed towards the fluidic channelmay be in direct contact with said channel. In other embodiments, thesurface directed towards the fluidic channel may have a dielectriccoating completely covering the surface of the sandwiched region whichwould otherwise be in direct contact with said fluidic channel. In otherembodiments, the surface directed towards the fluidic channel may have ametal coating completely covering the surface of the sandwiched regionwhich would otherwise be in direct contact with said fluidic channel. Inyet other embodiments, a metal layer and a dielectric layer maycompletely cover the surface of the sandwiched region which wouldotherwise be in direct contact with said fluidic channel.

The two of the other semiconductor regions, the regions accomplishingthe sandwiching of the aforementioned sandwiched region, are located atleast partially on the electrically-insulating layer. The two of theother semiconductor regions are doped at a higher concentration than thesandwiched region, and are electrically connected to the end electrodes.In response to the change in surface potential due to the targetmaterials in the fluidic channel, the at least three contiguoussemiconductor regions facilitate an interaction, largely using thesandwiched region proximate fluid in the fluidic channel, to sense thepresence of the target materials in the fluidic channel by passingcurrent in a conducting mode from one of the end electrodes to the otherof the end electrodes. The passing current is indicative of theconductance characteristic of the sandwiched semiconductor region,wherein said conductance may change in response to the presence of saidtarget materials. In certain specific embodiments, the conducting modeis a depletion mode. In other embodiments, the conducting mode is anaccumulation mode. The change in conductance in response to the targetmaterials may be an increase or a decrease in the conductance of thesandwiched semiconductor region.

The apparatus further may include an amplification circuit, which may besupported by the SOI structure or the substrate semiconductor, and thatis designed to facilitate sensing the target material near the fluidicchannel by way of a change in the electric field.

In certain embodiments, the apparatus also includes an alternative setof at least three contiguous semiconductor regions. The alternative setis electrically isolated from the aforementioned first set by a layer ofdielectric, and has a different doping specificity (e.g., higher orlower as compared to the first set of contiguous semiconductor regions).The alternative set may be arranged to operate in place of the firstmentioned at least three contiguous semiconductor regions, or may be inaddition to all three of the three contiguous semiconductor regions. Incertain embodiments, the surface of the sandwiched semiconductor region,or the surface of another layer which may cover or overlay thesandwiched semiconductor region, may be exposed toward the fluidicchannel, and may be textured to increase sensitivity.

The apparatus of the instant disclosure can be utilized in a number ofways. In an example embodiment, a fluid having target materials isprovided through the fluidic channel. An electric field is applied inand around the channel, which facilitates sensitivity by inducingmovement of molecules in the fluid. In certain specific embodiments, theelectric field utilizes electrophoresis in application. In otherembodiments, the electric field utilizes electro-osmosis. In addition toapplying an electric field, a bias voltage may be applied to thesandwiched region by a reference electrode or back bias source, and thepresence of the target materials is sensed by detecting a change in thecurrent passed in a conducting mode in response to the voltage inducedby the target materials, which may be further enhanced by an appliedbias voltage. In an example embodiment, the method of utilizing theapparatus includes transitioning from a low-conducting mode to thehigh-conducting mode in sensing the presence of target molecules or viceversa.

In an example method of manufacturing an apparatus discussed herein, theSOI device is formed over (e.g., on or in a portion of) the insulatinglayer. In certain embodiments, the insulating layer is a buried-oxidelayer. Further, the at least three semiconductor regions are doped tothe common polarity desired. In certain embodiments, a set of at leastthree alternative contiguous semiconductor regions are formed on theinsulating layer. The alternative set has a different doping specificityas compared to the first set of at least three semiconductor regions,and can operate in place of the first mentioned of the at least threecontiguous semiconductor regions.

Turning now to the figures, FIGS. 1A and 1B illustrate an exampleembodiment of an apparatus constructed in accordance with the instantdisclosure. As seen in the perspective views therein, the apparatusincludes an SOI structure. In the embodiment shown, the SOI structureincludes a silicon base 120, an insulative (e.g., oxide-based) layer 125and a fluidics structure 130. In other embodiments, the layer 125 can bea semiconductor substrate with the same or opposite polarity as thethree contiguous regions as above, or the SOI construction is usedwithout removing/etching the Silicon at all (as in the case where thesandwiched region is accessed on only one of the two opposing surfaces).The fluidics structure 130 may be fabricated of PDMS(“Polydimethylsiloxane”). In one embodiment a fluidic channel 100 isdefined by a portion of the upper-facing surface of the insulative layer125 and by a bottom-facing surface of the fluidics structure 130. Thesemiconductor device 110, as seen in FIG. 1B, illustrates a fluidsolution (inlet) 140 which may include target molecules to be sensed.

Shown in FIG. 2 is a manufacturing process flow according to an exampleembodiment of the instant disclosure. The process includes an SOI wafer.In certain embodiments, the silicon thickness is approximately 100 nm,and buried oxide (BOX) thickness of 400 nm. One or multiplesemiconductor devices are patterned onto the wafer using opticallithography. In certain embodiments, the length of each semiconductordevice is approximately 40 μm, and the width can be adjusted from 300 nmto 500 nm, in other embodiments, the length and width can be varied; forexample, the length may be from 40 μm to 100 μm or more, from 10 μm to40 μm, from 1 μm to 10 μm, or may be less than 1 μm, and the width maybe from 100 nm to 300 nm, from 30 nm to 100 nm, from 500 nm to 2 μm,from 2 μm to 10 μm, or more, or the width may less than 30 nm assemiconductor manufacturing processes improve. In certain specificembodiments, there are 20 identical, parallel semiconductor sensors ineach device. In other embodiments, there may be an array semiconductorsensors in each device. There may be 20 to 500, 500 to 5,000, 5,000 to50,000, 50,000 to 500,000, 500,000 to 5,000,000, 5,000,000 to50,000,000, or more than 50,000,000 sensors in each device. The nextstep is to etch off the silicon area around the semiconductor devicesdone by dry etching. This is followed by another lithography step todefine the place of the fluidic channel under the semiconductor devices.The semiconductor devices are also used as a mask for the wet etching ofthe buried oxide underneath them. Hence, a long channel with a depth of400 nm (thickness of buried oxide) is etched out. In certainembodiments, the regions at the end of the bridges are implanted withboron to make a p+ or P++ region for the ohmic contact to the aluminumelectrodes. Parallel to the fabrication process of the semiconductordevices, a mold for the fluidic channel is made. Liquid PDMS is pouredinto the mold and, once solidified, is peeled off from the moldsubstrate. The PDMS fluidic channel is then exposed to oxygen plasma,aligned to the wafer containing the bio-sensors and bonded to the wafer.In certain specific embodiments, the chip is left in a 70 C ovenovernight for better PDMS to silicon oxide adhesion.

FIG. 3 shows another example embodiment of the apparatus for sensingtarget materials. The apparatus shown in FIG. 3 includes an SOIstructure 380, and a fluidic channel 330 supported thereon. Alsoincluded is a semiconductor device 300. The semiconductor device 300 ofthis example embodiment includes three electrically-contiguoussemiconductor regions, which are doped to a common polarity. The threeelectrically-contiguous semiconductor regions include a sandwichedregion 360, and two outside other regions 340 and 350. As discussedabove, the regions 340 and 350 are doped to a higher concentration thanthe sandwiched region 360. As can be seen in FIG. 3, the sandwichedregion 360 has a surface directed toward the underneath fluidic channel330. The fluidic channel 330 passes target materials through thechannel, and target molecules may induce a voltage which change thesurface potential of semiconductor and the sandwiched region mayrespond, by changing the conductance of said semiconductor region.

Also shown in FIG. 3, on the SOI structure 380, an amplification circuitwhich may comprise at least a bipolar transistor, including a base 390,a collector 392, and an emitter 394 which facilitates sensing of thetarget material flowing through the fluidic channel 330 by way ofamplifying the current which passes through sandwiched semiconductorregion 360. In other embodiments MOSFET, JFET or any other circuitsknown in the art may be used for amplification or readout. A metalelectrode may be used to bring current into or out of SOI structure 380.FIG. 3 also shows the two other semiconductor regions 340 and 350 beingelectrically coupled (optionally) to electrodes 310 and 320 (discussedbelow in connection with FIG. 4). The electrodes 310 and 320 may be usedto generate electrophoretic flow or electroosmotic flow.

The semiconductor device 300 may be arranged to respond to a biasvoltage. The bias voltage may be applied to the target materials in thefluidic channel 330 by a reference electrode (not shown), or by backbias source (not shown), and said bias voltage will be added to theinduced voltage resulting from target materials whereby thesemiconductor device 300 senses the target materials in responsethereto.

In certain example embodiments, the surface of a semiconductor devicedirected toward the fluidic channel is roughened to increasesensitivity. Shown in FIG. 4A is a roughed version of the sandwichedsemiconductor region 400. In certain specific embodiments, the roughenedsurface 400 includes beads or particles for adding roughness or forusing as mask for etching and roughening the surface 410. Shown in FIG.4B is a simulated concentration of target solution before and afterpassing the roughened surface 400 and the non-roughened surface 420.

Shown in FIGS. 5A and 5B are examples of the electro-osmosis andelectrophoresis application. FIG. 5A shows application of an AC signalon two neighboring electrodes 510 and 520 to create an electric-fieldsolution flow 500. FIG. 5B shows a schematic of the electro-osmosis andthe electrophoresis application, positioned around the sensor tocirculate the flow and capture the negatively-charged target DNAmolecules at the same time. By changing the cycle of AC signal, thedirection of electro-osmosis force is not reversed and continuescirculation in the same direction as the other half cycle of the ACsignal.

FIG. 6A shows an example embodiment of semiconductor devices inaccordance with the instant disclosure. Three sets of three contiguousregions are shown in FIG. 6A as placed with respect to a fluidic channel670. A first sandwiched region 600 is adjacent the fluidic channel 670.The first sandwiched region 600 includes a surface that is directedtoward the fluidic channel 670 for sensing the target materials passedthrough the channel. Sandwiching the first sandwiched region 600 are afirst end region 610 and a second end region 640. The first sandwichedregion 600, the first end region 610, and the second end region 640 makeup a first set of three contiguous regions. The first sandwiched region600, the first end region 610, and the second end region 640 may beseparated from at least one additional set of three contiguous regionsby a dielectric layer 650. The additional sets of three contiguousregions include a first alternative end region 620, an alternativemiddle region 630, and a second alternative end region 660. Turning toFIG. 6B of an additional embodiment of the instant disclosure, asemiconductor device contains three contiguous regions. The contiguousregions include the sandwich region 680, which includes a comb-likestructure, and connects the two other contiguous regions 690 and 695.The three contiguous regions of both embodiments shown in FIGS. 6A and6B are doped to a common polarity. In certain example embodiments, thesandwich regions 600, 630, and 680 are doped to have a lower-dopantconcentration than their respective other (outside sandwiching) regions610/640, 620/660, and 690/695.

Various embodiments are directed to semiconducting nanobridge (or“nanowire”) types of sensors, such as those including silicon, siliconof a silicon-on-insulator (SOI) structure, or other semiconductormaterial.

Other example embodiments are directed to the operation of a biosensordevice. Phases of operation include: a loading phase in which probebiomolecules are attached or associated with a sensor, and a detectionphase in which one or more target biomolecules, reaction products, orreaction byproducts have changed the local voltage generated orimpressed on the nanobridge. The conductance of the nanobridge/nanowireis modulated during the loading and detection phases.

Aspects of the present invention are directed to modeling operationalcharacteristics of a biosensor, and setting the characteristicsaccording to the various modeling results. In many implementations, abio-linker molecule such as APTES (10%), or ssDNA (1 μm), or acombination thereof, which may be used in combination with dsDNA, isused in a layer on a sensor device surface to achieve an effective orotherwise desirable charge and related threshold voltage shift, relativeto the sensitivity of device to molecules such as dsDNA. Such anapproach may involve, for example, using one or more aspects asdiscussed in T. Sakata, et al., Japanese Journal of Applied Physics,Vol. 44, No. 4B, pp. 2854-2859, 2005, which is fully incorporated hereinby reference.

In various embodiments, a back bias voltage is used to set or otherwiseinfluence the detection sensitivity of such a sensor, and can further becontrolled to establish a condition sensitive to a particular targetspecies, such as for use with a p-type silicon device under positiveback bias voltage. One such implementation is shown in FIG. 3 ofAppendix A filed in the underlying provisional application, showing aneffect of an example back bias voltage application upon devicesensitivity for a 10e13 (1/cc) P-type substrate. In various contexts,the effect of a back bias voltage in bringing the device to subthresholdregion in SOI devices and increasing device sensitivity is stronger thanin planar MOSFET devices.

Various embodiments are directed to using back bias as well as a chargedbiomolecule layer to modulate the Fermi level energy in an active areaof device. This approach can be used to change (increase) the slope ofId-Vg curve, which is equivalent to sensitivity. Using this approach,the device can be operated in a desirable (e.g., sensitive) regime.Moreover, this regime can be tailored to sense specific biomolecules.

Various aspects of the present invention, as directed to the applicationof a bias via one or more gates or reference electrodes and asapplicable to a back bias, may be implemented as follows. A positivecharge associated with biomolecules is used to deplete hole carriers ina p-type sensor and increase the barrier height in the sensor channel,with a resulting drop in conductance. A negative charge associated withDNA may increase conductance. A back bias can be used in this context tocontrol the conductance modulation. Alternatively, a bias may be appliedby a combination of a reference electrode and a back bias.

The various biosensors and biosensor devices as discussed herein may befabricated using a variety of approaches. One such device is shown inFIG. 5 in Appendix A filed in the underlying provisional application,including an array of biosensors fabricated from a SOI wafer (e.g., withsuch a wafer having a top silicon thickness of about 50 nm).

For sensing DNA, different linkage molecules are used for attachment tosensor surfaces to achieve a desired charge density per unit arearelative to each linkage molecule used and/or desired sensitivity.Various embodiments employ a nanobridge geometry (e.g., as shown in FIG.1 of Appendix A), used as a planar sensor surface for detection and highsensitivity, via double-side exposure to target molecules such as DNAand/or protein. Planar biosensors can be manufactured with a relativelylarge width and be implemented with a relatively large signal current,achieving desirable noise reduction and a high signal-to-noise ratio(SNR). Various embodiments as shown in the Appendices describeconductance modulation via sensor functionalization, such as by DNAlinkage molecules. Current-voltage (I-V) characteristics for suchdevices are shown, for respective embodiments involving specificconcentrations of linkage biomolecules, including APTES, Strepavidin andPoly- L-Lysine (PLL), which have positive charge, such as shown in FIG.7 of Appendix A, filed in the underlying provisional application. Forthe APTES molecules, an example solution includes 2% APTES in Ethanol orAceton; for the Streptavidin molecules, an example solution includes 0.1mg/mL streptavidin in PBS with 2 mM biotin solution. Such bio-speciesmay be obtained, for example, from Sigma Aldrich Company of St. Louis,Mo.

Other example embodiments are directed to on-chip amplification of asignal produced via a biosensor and approach as discussed herein, toimprove the signal-to-noise ratio. The detection signal, correspondingto conductance characteristics of a channel type region of a biosensorsuch as the previously described nanobridge as coupled to targetmolecules, is amplified. The noise can also be averaged as well. Suchapproaches may employ, for example, a Darlington or single amplifier.Other embodiments may utilize a more sophisticated amplifier and readoutcircuit, wherein row and column selection may be included as part of theamplifier circuit. The operational mode of the device may be implementedin the fully depleted, inverted, accumulated, enhanced or resistive modedepending on the concentration and operational circumstances, or the useof a BJT or MOS amplifier, such as discussed in one or more of theAppendices that form part of this document.

Example types of biosensor systems, as may be implemented in accordancewith various embodiments, include a bio-species coated bead system and adirect binding system. In the direct binding system, the operation ofthe device has two main phases: a loading phase in which probebiomolecules (e.g., ss-DNA) are attached or associated with the sensor,and a detection phase in which a target bio-molecule (e.g., nucleotides)is attached or associated with the probe molecule. The conductance ofthe nanobridge/nanowire is modulated during the loading and detectionphases.

Using the bead system approach, biomolecules (e.g., DNA or proteins) areattached or associated with a bead. The beads are then injected into thesystem and allocated adjacent to the biosensor. This allocation mayinvolve, for example, manipulation of the flow of beads so as tointroduce the bio-species molecules to the sensor. Various bimolecularinteractions and reactions may then be monitored.

Accordingly, various example embodiments are directed to a sensor devicehaving a planar array with on-chip amplification, as discussed herein.Various aspects are directed to large scale arrays of such sensors,which can be tailored to specific applications and set, or tuned, viathe application of a bias. Such aspects can be used to achieve adesirable SNR and dynamic range, and are suitable for DNA sequencing andthe detection of DNA, protein or other bio-species in microfluidicstructures. Various embodiments use rough surface structure features toincrease the sensitivity and SNR (e.g., as described in one or more ofthe Appendices forming part of this document).

For general information regarding sensors, and for more specificinformation regarding biosensors and related applications as may beimplemented in connection with one or more example embodiments describedherein, reference may be made to the following references, each of whichis fully incorporated herein by reference: H. Esfandyarpour,NSTI-nanotech, Vol. 4, p. 421, 2007; J. Fritz, et al., PNAS, Vol. 99,No. 22, p. 14143, 2002; and C. M. Lieber, PNAS, Vol. 101, No. 39, p.14017, 2004.

Experimental Embodiments

For any electrical device to be able to detect a charge, the chargeshould bind or associate to or near the device's surface and affect theelectrical characteristics of the device.

DNA molecules do not generally bind to oxide surfaces on their own. Forthis reason, linker molecules may be provided for their attachmentthereto. These linker molecules could mediate DNA binding to the siliconoxide surface through physical or chemical forces. For instance, PLL,APTES and Streptavidin-Biotin are a physical (ionic) binding linker,while pHPMA binding to a surface may be a chemical (covalent) bond.

After functionalizing the surface of the sensor with the linkermolecule, there will be a shift in the electrical characteristics of thesensor due to the induced charge of the linker layer on the surface ofdevice and associated changes in the counter ions in the fluid. Thefirst layer of DNA, called probe DNA, will induce another change (shift)in the electrical characteristics and counter ion electricalcharacteristics. A solution containing the sample DNA is then introducedinto the system using simple fluidics and it is tested forcomplementarity to the probe DNA. If the sample contains thecomplementary strand, it will hybridize to the probe DNA attached to thesensor and induce a secondary shift in the electrical response curve,and the thus the conductivity of the nanosensor. Usually this secondshift is smaller than the first one, which could be due to the fact thatdouble-stranded DNA is rigid in contrast to single-stranded DNA which ismore flexible. The rigidity of the double-stranded DNA will result in agreater average physical distance between the DNA and the surface of thesensor and hence it will lower the effect on the change in conductivityof the device (see A. Poghossian, A. Cherstvy, S. Ingebrandt, A.Offenhausser, M. J. Schoning, Possibilities And Limitations OfLabel-Free Detection Of DNA Hybridization With Field-Effect-BasedDevices, Sensors and Actuators B 111-112,470-480 (2005)). In addition tothe rigidity of double-stranded DNA, there is always a possibility thatthere is interference with the binding of the sample DNA to the probeDNA because of the already highly functionalized and charge-modifiedsurface of the sensor. This could be a result of the coulomb repulsionbetween two negatively-charged DNA molecules. For any electrical deviceto be able to detect a charge, charge should bind or associate to ornear the surface of the device. The negative charge of bound DNAmolecules will affect current or conductance of a device.

An electrical sensor should have a high surface-to-volume ratio in orderto have the highest possible signal in response to a charge modulationon or near the surface. The sensor should have an easily measurablesignal-to-noise ratio because the change in electrical signal due to thecharge modulation is usually only a few percent of a base-line referencecurrent. For best signal to noise ratio, in addition to the need for alarge surface-to-volume ratio, the active area of the sensor should alsobe as large as possible to capture as many DNA molecules or other targetmoieties as possible.

In one embodiment, a nanowire sensor as described herein has a rodstructure with high aspect ratio. The nanowire may be comprised of adoped semiconductor and can be formed through chemical vapor deposition,lithography or other methods known in the art. The nanowire sensordevice may have three elements. Two conductive formed regions and theinsulator or low conductivity gap between the two regions. Theconductive regions can be made of highly doped semiconductor or metal,through photolithography.

There are two designs that could be proposed for such a sensor: an arrayof nanowires; and an array of parallel suspended plates. The firstdesign consists of an array of vertical or horizontal nanowires (withthe sensor structure along a largely horizontal plane), in which thedirection of the solution flow is perpendicular to the nanowires. DNAmolecules or other moieties will bind or associate to or near thefunctionalized nanowires and change the conductance due to their charge.In the second design, an array of parallel nanobridges makes up thesensor. As solution flows through the channel, DNA molecules or othermoieties will bind or associate to or near both sides of the bridges,top and bottom, similar to a double gated device. The charge of the DNAmolecules or other moieties will cause accumulation or depletion of moreholes in the bulk of a p-type sensor and hence increase its conductance.

Conductance of the sensor will also change after functionalizing thesensor surface. This change depends on the charge of the linker moleculelayer. If negative, the bulk of the p-type sensor will accumulate holesand conductance will increase. On the other hand, if positive,conductance will decrease due to the hole depletion. Linker moleculesusually have a positive charge, especially when they are binding DNAcovalently, therefore the conductance of the sensor will decrease at thetime of functionalization.

The sensor is designed to be sensitive enough to detect lowconcentrations of DNA and at the same time highly effective at capturingthe sample or target DNA molecules in order to detect binding. Tocompare which design is better for capturing the DNA molecules, thetotal number of DNA molecules bound to the sensors was compared usingCOMSOL simulation. In this simulation, the nanowire and nanobridgedesigns are compared with each other. Both sensors have the same totalsurface area and the same surface-to-volume ratio.

The reaction at the active surface is given as in Equation 2.1:

k_(ads) c + θ ⇔ c_(s) k_(des)Where c is the bulk concentration, θ is the surface concentration ofactive sites, c_(s) is the surface concentration of adsorbed species(moles per unit surface), kads is the rate constant for the forwardreaction, and kdes is the rate constant for the backward reaction.

This simulation assumes the following values regarding the hybridizationkinetics of DNA on surface and the diffusion rate on surface andsolution (see, e.g., Hong Shen et al., DNA Diffusion in Mucus: Effect ofSize, Topology of DNAs, and Transfection Reagents, Biophysical JournalVolume 91 July 2006 639-644; and J. R. Pascault, A Finite Element Studyof the DNA Hybridization Kinetics on the Surface of MicrofluidicDevices, A thesis submitted to the faculty of the Worcester PolytechnicInstitute):

-   -   k_(des)=6×E-5 s-1    -   k_(ads)=6×E4 M-1 s-1    -   Kd=k_(ads)/k_(des)=1E9 M-1    -   Convection & diffusion for particles movement    -   Vo=1E-3 m/s    -   D=1E-9 m2/s    -   Ds=0    -   Bo=1E-3 moles/m2

kads: the rate constant for the backward reaction,

kdes: the rate constant for the forward reaction,

Vo: the velocity of incoming fluid,

D: the diffusion rate of DNA molecules thorough the solution,

Ds: the diffusion rate of DNA molecules on the surface of sensor,

Bo: the density of bonded DNA molecules on the surface.

For having equilibrium in the adsorbed material, cs, on the surface ofthe sensor device, the equation governing the surface diffusion and thesurface reaction rate is shown in Equation 2.2:

${\frac{\partial c_{s}}{\partial t} + {\nabla{\cdot \left( {{- D_{s}}{\nabla c_{s}}} \right)}}} = {{k_{ads}c\;\theta} - {k_{des}c_{s}}}$where Ds is surface diffusivity. However, the concentration of activesites is equal to the difference between the total number of activesites and the number of sites that are occupied by the adsorbedmolecules, therefore, the equation for the rate reaction is given byEquation 2.3:

${\frac{\partial c_{s}}{\partial t} + {\nabla{\cdot \left( {{- D_{s}}{\nabla c_{s}}} \right)}}} = {{k_{ads}{c\left( {\theta_{0} - c_{s}} \right)}} - {k_{des}c_{s}}}$where θ₀ is the total number of active sites available on the surface ofthe sensor. The initial condition of this equation is that theconcentration of the adsorbed species on the surface of the sensor iszero at the beginning of the process (c_(s)(t=0)=0). The equation forthe surface reaction includes the concentration of the bulk species, c,at the top of the sensor surface. The equation must be solved for thesurface reaction in combination with the mass balance in the bulk. Thetransport in the bulk of the channel is described by aconvection-diffusion equation, Equation 2.4:

${\frac{\partial c}{\partial t} + {\nabla{\cdot \left( {{{- D}{\nabla c}} + {cu}} \right)}}} = 0$where c is molecule concentration at the channel, D is the diffusivityof those molecules in the channel, and u is the velocity vector.

The initial condition for c is set so that the bulk concentration isequal to co at the beginning of process (c(t=0)=c₀). The adiabaticcondition for all the surfaces except the inlet, outlet and the sensorsurface for the material balance is:n·(−D _(s) ∇c _(s))=0

The boundary condition at the sensor surface couples the rate of thereaction at the surface with the flux of the binding molecules, theconcentration of the adsorbed molecules and the concentration of themolecules in the bulk:n·(−D∇c+cu)=−k _(ads) c(θ₀ −c _(s))+k _(des) c _(s)

The boundary conditions for the inlet and outlet are:

-   -   Inlet: c=c₀    -   Outlet: n·(−D∇c+cu)=n·cu

The boundary condition for the other surfaces except for the sensor is(adiabatic):n·(−D∇c+cu)=0

The number of DNA molecules bonded to the planar sensors wasapproximately 1.6 E-15 (moles/cm²), and approximately 3.71 E-16(moles/cm²) for the nanowire sensor (see, e.g., Manual of COMSOLMultiphysics). Nanowire sensors captured less DNA molecules due to theirgeometry, as the DNA-depleted solution passing the first rows innanowire array hits the rest of nanowires and decreased the number ofpossible capturable DNA molecules.

The following discusses various design types of bio-sensor devices. Onetype can be classified as a bulk MOS device, based on p-type siliconconstruction and with an electrical connection through a highly p-typeregion to two ohmic contacts. A second type is a single-sided SOI MOSdevice that keeps the sandwiched semiconductor region intact but has avery thin layer of silicon below the said sandwiched semiconductorregion. A third double-sided SOI type has an fluidic access to bothsides of sandwiched semiconductor region. In other embodiments the MOS(metal oxide semiconductor) structure can be configured without metallayer or in other embodiments without this insulating layer or acombination of the two.

Electrical characteristics of these three designs was simulated byapplying a voltage bias to one of the two end semiconductor regions andanother sweeping voltage to the sandwiched semiconductor region, andmeasuring the passing current at the same time by an MEDICI simulator.As can be seen from the curves, shown in FIG. 7, the first design (bulkMOS sensor) exhibits the highest level of current due to the thick bulkbut very shallow slope for the Id/Vg characteristics. The third design,i.e., double sided structure, has the steepest slope and the highestsecond derivative of the Id/Vg characteristics. A steeper slope meansthat by modulating the charge on the sandwiched semiconductor region(DNA charge), there will be a larger current change (higher sensitivityto charge modulation) and the higher second derivative means that theratio of the current change in response to target DNA hybridization andthe probe DNA immobilization is higher (more change in current afterhybridization).

The effective charge of DNA on the sensor surface after probeimmobilization and target hybridization can be calculated from theexperimental data from others' electrical DNA detection experimentresults, as is well known.

By importing this data into the MEDICI device simulator, the ratio ofconductance change in target hybridization and probe immobilization canbe found for these three devices. The double gated SOI MOS sensor hasthe highest ratio for the conductance change. It is proposed that asensor is positioned on an etched channel in the substrate and there isa PDMS micro-fluidic device with a fluidic channel formed therein on topof the sensor. Therefore, the sensor is suspended between two fluidicchannels (one on top and one at the bottom) and DNA molecules can bindto both sides of the thin sensor, making it like a double gated SOI MOSsensor.

The bio-sensor should be able to detect an ionic solution flowingthrough the channel when the top micro-fluidic channel is bonded to thesensor substrate. When the pH of the solution increases from a neutralor low pH, hydroxide ions which are bound or associated with the surfaceof the sensor accumulate holes near the p-type sensor surface. Thisleads to an increase in the number of carriers and consequently anincrease in conductance. On the other hand, by decreasing the pH of thesolution, a reduction in the number of hydrogen ions bound or associatedwith the surface of the sensor will deplete holes near the p-type sensorsurface so that the conductance of the sensor decreases.

Experimental tests of the sensor with different pH solutions anddifferent functionalizing linker molecule showed that by increasing thepH of the solution, there is an increase in the conductance. When thesensor is dry, its current is around 40 uA. By flowing a PBS bufferhaving a pH of 7, there is an increase in conductance. This increase isdue to the difference between the work function of air and the workfunction of the solution that causes an accumulation of holes in thep-type sensor surface, and or as a result of binding of hydrogen ions tothe surface of the sensor.

When the device surface is functionalized with APTES, a physical linkermolecule, conductance of the sensor decreases because of the positivecharge of the APTES layer that depletes the holes from sensor surface.On the other hand, Streptavidin, a slightly negatively-charged linkermolecule, slightly increases the current of the sensor due to the chargeof the linker layer. Streptavidin has almost no effect on the current ofthe sensor when it flows through the channel with an APTES coatedsensors. This implies that the charge of Streptavidin is almostnegligible in comparison to the charge of the APTES layer. However,there is also the possibility that the layer of Streptavidin is notclose enough to the surface of the sensor to be able to affect itscharge distribution because the APTES already coats the surface of thesensor with around 3-8 mono-layers. The distance that a charged moleculeneeds to be within to cause a significant effect on the sensor is afunction of the Debye length, and thus of the ionic concentration of thefluid in the volume adjacent the sensor.

In order to detect DNA hybridization, the sensor surface should befunctionalized with immobilized probe DNA. By flowing the target DNAsolution through the channel, the device will detect whether the DNAcomplementary or non-complementary. When two strands of DNA arecomplementary, they will react and bind to each other. This binding willincrease the negative charge on the surface, because there are twostrands of DNAs on the surface instead of one. This increase in negativecharge increases the conductance of sensor by accumulating more holesnear the surface of the sensor. When two strands are not complementary,they do not bind to each other and the charge on the surface will notchange. However, there is usually an increase in conductance even whenthe target DNA is not complementary to the probe DNA. This is due to thefunctionalizing layer, which is positively-charged. The charge of thelinker molecule layer compensates the negatively-chargednon-complementary DNA through coulomb interaction. Hence, an increase incurrent is detected.

The sensitivity of the bio-sensor can be increased by enhancing thenumber of captured molecules or by enhancing the signal-to-noise ratio.

A bipolar junction transistor (BJT) can be used as AC or DC amplifier. ABJT is a three terminal device with a “base”, “emitter” and “collector”made of doped semiconductor (e.g., NPN bipolar transistor). Thedirection of current flow from the sensor is always between the baseterminal and emitter terminal. Both electrons and holes transports areinvolved in the operation of bipolar junction transistors. Electricalcurrent is generated by diffusion of charge carriers from one region tothe other two regions with different charge concentration.

In active mode operation, the base-emitter junction diode is forwardbiased and the base-collector junction diode is reverse biased. In anNPN transistor, for example, when a positive voltage is applied to thebase-emitter junction, the equilibrium between thermally generatedcarriers and the repelling electric field of the depletion regionbecomes unbalanced, allowing thermally excited electrons to inject intothe base region. These electrons diffuse through the base from theregion of high concentration near the emitter toward the region of lowconcentration near the collector. In the collector and the emitter,holes are minority carriers because they are n-type doped, so that theelectrons are majority carriers. On the other hand, the electrons in thebase are minority carriers because the base is p-type doped.

To minimize the recombination of carriers that are crossing the basebefore reaching the base-collector junction, the transistor's baseregion must be thin enough that carriers can diffuse across in a timeless than the minority carrier's lifetime in the base, and the basethickness must be less than the diffusion length of the minoritycarriers in the base (e.g., electrons in a NPN bi-polar junctiontransistor). The collector-base junction diode is reverse biased inactive mode, and a small number of electrons come from the collector tothe base, but electrons that diffuse from the emitter (and generatedones in base) to the base towards the collector are swept into thecollector by the electric field in the depletion region of thecollector-base junction. The currents in BJT terminals can be controlledso that they are exponentially-dependent on the base-emitter voltage.Since the base-collector junction diode is reverse biased in the activemode operation, the base-collector voltage does not have much of aneffect on the emitter-collector current. The common emitter currentgain, β_(F), is approximately the ratio of the DC collector current tothe DC base current in the forward-active region. The current gain,β_(F), is given by:

$\beta_{F} = \frac{I_{C}}{I_{B}}$

For the fabricated BJT, this ratio was around 30. By increasing the basevoltage, collector current increases, and by binding DNA to the surfaceof the sensor, collector current again increases due to accumulation ofholes in the sensor surface region. When conductance of the sensor inseries with the base increases, the current passing through thecollector increases as well.

If lower concentrations of DNA are used, there will be fewer DNAmolecules bound to the sensor. In order to increase the number of DNAmolecules bound to the sensor, thereby improving the sensor signal, moreDNA can be interact with the DNA bound or associated with the surface ofthe sensor by dielectrophoresis and electro-osmosis forces.

Electrophoresis is the motion of dispersed particles relative to a fluidunder the influence of an electric field The dispersed particles have anelectric surface charge, on which an external electric field exerts anelectrostatic coulomb force. According to the double layer theory, allsurface charges in fluids are screened by a diffuse layer of ions, whichhas the same absolute charge but opposite sign with respect to that ofthe surface charge. The electric field also exerts a force on the ionsin the diffuse layer, that force having a direction opposite to theforce acting on the surface charge. The force of the electric field isapplied to both the particle and to the ions in the diffuse layer thatare located at some distance from the particle surface. Through viscousstress, part of the electric force is transferred to the particlesurface. This part of the force is also called electrophoreticretardation force. The electrophoretic retardation force may be greaterthan or less than the electrostatic force which also acts on theparticle.

Electroosmotic flow is the motion of liquid induced by an appliedpotential across a material, capillary tube, membrane, channel (ormicrochannel), or any other fluid conduit. Because electroosmoticvelocities are independent of conduit size, as long as the double layeris much smaller than the characteristic diameter scale of the channel,electroosmotic flow is most significant in small channels.Electroosmotic flow is an essential component in chemical separationtechniques. Electroosmotic flow is caused by the Coulomb force inducedby an electric field on a net mobile electric charge in a double layer.Because the chemical equilibrium between a solid surface and anelectrolyte solution typically leads to the interface acquiring a netfixed electrical charge, a layer of mobile ions, known as an electricaldouble layer or Debye layer, forms in the region near the interface.When an electric field is applied to the fluid (usually via electrodesplaced at inlets and outlets), the net movement of charge in theelectrical double layer is induced by the resulting Coulomb field. Theresulting flow is termed electro-osmotic flow.

DNA concentrators using AC electroosmosis have been developed toconcentrate large molecular weight molecules. Hence, electrophoresis,which is the most well-known electro-kinetic phenomenon, can provide anattractive force to hold small charged particles.

AC electro-osmosis can generate bulk fluid flow to transportsingle-strand DNA molecules from a large effective volume to theelectrode surface. Electrophoresis can attract single-strand DNAmolecules and hold them on the electrode surface simultaneously. When anelectric field generated by the two electrodes is applied tangentiallyto a surface bathed in electrolytes, the charges in the electricaldouble layer experience a force. Consequently, the fluid is pulled alongthe charges and bulk fluid flow is generated. Electric potential causescharges to accumulate on the electrode surface, creating a chargedensity that forms the electrical double layer is formed. The doublelayer interacts with the tangential component of the electric field toinduce fluid motion along the electrode surface. AC electro-osmotic flowcan be induced by an AC electric field on microelectrodes in thefrequency ranges below the charge relaxation frequency. At arelatively-low frequency, most of the potential drop occurs in thedouble layer. The electric field is at a minimum and a small slipvelocity is generated along the surface. At a high frequency, thecharges in the double layer are less and slip velocity tends to be zero.Accordingly, the maximum slip velocity occurs at an intermediatefrequency. For further information in this regard, reference may be madeto: A. Ramos, H. Morgan, N. G. Green, A. Castellanos, ACelectric-field-induced fluid flow in microelectrodes; J ColloidInterface Sci 217:420-422,1999; M. R. Bown A E C. D. Meinhart, ACelectroosmotic flow in a DNA concentrator, Microfluid Nanofluid (2006)2: 513-523; D. Stein, Z. Deurvorst, F. H. J. van der Heyden, W. J. A.Koopmans, A. Gabel and C. Dekker, Electrokinetic Concentration of DNAPolymers in Nanofluidic Channels, Nano Lett. 2010, 10, 765-772; H.ChengKin, F. Lei, K. Ying Choy, L. M. C. Chow, Single-stranded DNAconcentration by electrokinetic forces, J. Micro/Nanolith. MEMS MOEMS8_2_, 021107, 2009; K. Fong Lei , H. Chenga, K. Ying Choyb, L. M. C.Chowb, Electrokinetic DNA Concentration in Microsystems, Sensors andActuators A 156 (2009) 381-387.

Simulation of generation of electro-osmotic flow has been done for threeparallel electrodes in an ionic solution. The center electrode has adifferent applied voltage from the other two, and the other twoelectrodes are short-circuited to each other and have the same voltagebias. Some assumptions that have been used for the simulation are:

ρ 1000 kg/m3 Density of the fluid η 10−3 Pa · s Dynamic viscosity of thefluid ε_(r) 80.2 Relative electric permittivity of the fluid ζ −0.1 VZeta potential on the wall-fluid boundary σ 0.11845 S/m Conductivity ofthe solution D 10−11 m²/s Diffusion coefficient Vac 4 V AC voltage W1000 Hz frequency

The Navier-Stokes equations for incompressible flow describe the flow inthe channels:

${{\rho\frac{\partial u}{\partial t}} - {\nabla{\cdot {\eta\left( {{\nabla u} + \left( {\nabla u} \right)^{T}} \right)}}} + {\rho\;{u \cdot {\nabla u}}} + {\nabla p}} = 0$∇⋅u = 0

Here η represents the dynamic viscosity (kg/(m·s)), u is the velocity(m/s), ρ equals the fluid density (kg/m3), and p refers to the pressure(Pa). It is assumed that the flow has a fully developed laminar profileat the inlet channel. It is also assumed that at the outlet boundaryfluid flows out freely and the stress components at this boundary arezero.n·[−pI+η(∇u+(∇u)^(T))]=0

Most of the solid surfaces in contact with an electrolyte form a surfacecharge, and a double layer charge forms at the electrolyte/solid surfaceinterface in response to the spontaneously formed surface charge. Theelectric double layer forms because of the surface charges in thesolution/solid interface. By applying a voltage, the electric fieldgenerates the electro-osmotic flow that displaces the charged liquid inthe electric double layer. The electric force on the positively-chargedsolution close to the interface makes the fluid start to flow in thedirection of the electric field.

The thin electric double layer was replaced by theHelmholtz-Smoluchowski relation between the electro-osmotic velocity andthe tangential component of the electric field in all boundaries exceptfor the inlet and outlet:

$u = {\frac{ɛ_{w}\zeta_{0}}{\eta}{\nabla_{T}V}}$Where ϵ_(w) represents the fluid's electric permittivity (F/m), ζ₀ isthe zeta potential at the channel wall (V), and Vis the appliedpotential (V).

The balance equation for current density is expressed by Ohm's law. Ifthere is no gradient in the concentration of ions carrying the current,the divergence of the current density is set to zero:∇·(−σ∇V)=0where σ is conductivity (S/m). The electric potentials on the centerelectrode are sinusoidal in time with the maximum value of 4 V and thefrequency of 800 Hz. The adiabatic boundary condition on all boundariesexcept for the electrodes is that the normal component of electric fieldis equal to zero:−σ∇V·n=0The concentration of DNA molecules is increased after applying the DCand AC signal to capture them.

There are techniques to chemically bind the DNA to a specific surface.To immobilize probe DNA to a silicon oxide surface, a layer of APTES isoften used. APTES is known for covalently binding to silicon oxide butnot to silicon nitride. In order to be able to bind DNA molecules to thesensor's surface but not to other areas of the channel, silicon oxide isused for the gate dielectric of the sensor while all other surfaces inthe micro-fluidic channel are covered by a layer of silicon nitridewhich was deposited by a chemical vapor deposition technique with athickness of approximately 100 nm.

By flowing a solution containing 10% of APTES in ethanol, APTESmolecules covalently bound to the silicon oxide but not to the siliconnitride. All surfaces then were washed with ethanol to remove anyresidue and unbound APTES molecules in the channel. Probe DNA with aconcentration of 1 uM in PBS was then introduced into the channel. Inaqueous solutions, silicon nitride reacts with water and forms silica,silanol and ammonium groups. The ammonium dissolves in water andincreases the local pH of the solution from pH 6-7 to pH 9-10. Siliconnitride surfaces are negatively-charged in alkaline solutions (pH 9-10)since the isoelectric point of silicon nitride is between 5 and 6.Negatively-charged probe DNA molecules are covalently bound to the APTEScoated silicon oxide surface,. After the functionalization step, themicro-fluidic channel is washed with PBS buffer. DNA molecules in thisstep are labeled with EDC(N-(3-Dimethylaminopropyl)-N′-ethylcarbodiimide hydrochloride). Theprocedure for modifying 5′ phosphate groups to enable covalent bindingis:

-   1. Dissolve ethylenediamine (or alternative) to a final    concentration of 0.25 M in 10 μl of 0.1 M imidazole.-   2. Weigh 1.25 mg (6.52 μmol) of EDC into a micro-centrifuge tube.-   3. Add 7.5 μl of the prepared oligonucleotide to the tube containing    the EDC and immediately add 5 μl of the ethylenediamine/imidazole    solution.-   4. Vortex tubes until contents are completely dissolved, and then    briefly centrifuge the tube to gather contents.-   5. Add an additional 20 μl of 0.1 M imidazole, pH 6.-   6. Incubate reaction at room temperature for over night-   7. Remove non-reacted EDC and the by-products, and imidazole by    dialysis (e.g., Slide-A-Lyzer® MINI Dialysis Units) or spin    desalting column (Zeba™ Desalt Spin Column) using 10 mM sodium    phosphate, 0.15 M NaCl, 10 mM EDTA, pH 7.2, or other suitable    buffer.

solution of 10 μM BSA in PBS buffer is then flowed through the channel.Since BSA is a very sticky molecule, it binds to almost any cleansurface. In the channel, BSA molecules will bind to silicon nitride butnot to silicon oxide surfaces, which are already coated and saturatedwith APETS and the probe DNA molecules. The channel is then washed withPBS buffer to remove BSA molecules that did not bind to any surface.

Low concentration target DNA molecule solution for detection is flowedinto the micro-fluidic channel. Target DNA molecules will bind to theprobe DNA molecules when they are complementary. Since probe DNAmolecules are already bound to the bio-sensor surface, target DNAmolecules will also bind to the surface through the complementary DNAbound to said surface. DNA molecules are mainly bound selectively tosilicon oxide rather that silicon nitride, as the BSA largely preventsnonspecific binding to the silicon nitride surfaces.

Another way to enhance the number of DNA molecules bound to the sensorsurface is by increasing its effective area. This can be done byincreasing surface roughness so that the magnitude of roughness is onthe order of the DNA molecule size. A proposed way to achieve surfaceroughness on the order of 10 nm-20 nm is by etching silicon on a goldnanoparticle mask. A silicon surface could be etched by a KOH solutionat room temperature. The etch rate of silicon by cold KOH is around 10nm/min (see, e.g., K. R. Williams and R. S. Muller, Etch Rates forMicromachining Processing, Journal Of Microelectromechanical Systems,Vol. 5, No. 4, December 1996)).

In order to compare the capture rates in rough and smooth surfaces,COMSOL simulation was used. The assumptions behind these simulationsare:

-   -   Kads=6×10E4 M-1 s-1    -   Kd=Kads/Kdes=10E9 M-1    -   Kdes=6×10E-5 s-1    -   Vo=10E-3 m/s    -   D=10E-9 m²/s    -   Ds=0    -   Bo=10E-3 moles/m²

Kads: the rate constant for the backward reaction; Kdes: the rateconstant for the forward reaction; Vo: the velocity of incoming fluid;D: the diffusion rate of DNA molecules thorough the solution; Ds: thediffusion rate of DNA molecules on the surface of sensor; Bo: thedensity of bonded DNA molecules on the surface. The main assumptions forthese simulations are convection, diffusion and chemical adsorption, asdescribed by Longmuir equation.∝b/∝t=k _(on) c _(s)(b _(m) −b)−k _(off) b

The relationship of binding density and location on the sensor surfaceat different times relates to the integral of the binding curves as afunction of time and location, which yields the total number of boundDNA molecules. After 10 minutes, the smooth surface bound 1.6 E-16(moles/cm²) and the rough surface bound 2.06 E-10 (moles/cm²), whichcorresponds to around 25% more DNA molecules than on the smooth sensor.

Various embodiments described above, characterized in the claims and/orshown in the figures may be implemented alone, together and/or in othermanners. One or more of the items depicted in the drawings/figures canalso be implemented in a more separated or integrated manner, or removedand/or rendered as inoperable in certain cases, as is useful inaccordance with particular applications. For example, the variousdiscussions of field-effect transistors and other devices may beimplemented with different field-effect devices, using approaches asdescribed herein. In view of the description herein, those skilled inthe art will recognize that many changes may be made thereto withoutdeparting from the spirit and scope of the present invention.

What is claimed is:
 1. A method of manufacturing an apparatus forsensing a target material, the method comprising: patterning at leastone semiconductor device onto a silicon substrate, wherein the at leastone semiconductor device includes at least three electrically-contiguoussemiconductor regions and end electrodes; doping the at least threeelectrically-contiguous semiconductor regions to exhibit a commonpolarity, wherein the at least three electrically-contiguoussemiconductor regions include a sandwich region sandwiched between twoof the at least three electrically-contiguous semiconductor regions, andwherein the two of the at least three electrically-contiguoussemiconductor regions are doped at a higher concentration than thesandwiched region and are electrically connected to the end electrodes;and defining a location, on the silicon substrate, for a fluidic channelunder the at least one semiconductor device, wherein the at least threeelectrically-contiguous semiconductor regions are configured andarranged adjacent to the fluidic channel with a surface directed towardthe fluidic channel for coupling to the target material in a fluid inthe fluidic channel, and for responding to a change in potentialattributable to the target material.
 2. The method of claim 1, whereindefining a location for a fluidic channel under the at least onesemiconductor device includes etching a portion of the silicon substratelocated under the at least one semiconductor device to define thelocation for the fluidic channel.
 3. The method of claim 2, wherein thesilicon substrate is a semiconductor-on-insulator (SOI) structureincluding an electrically-insulating layer and wherein etching theportion includes etching a channel within the electrically-insulatinglayer with a depth of a thickness of buried oxide.
 4. The method ofclaim 2, wherein etching includes performing a lithography process. 5.The method of claim 1, further including etching a silicon area aroundthe at least one semiconductor device.
 6. The method of claim 5, whereinetching the silicon area includes performing a dry etch process.
 7. Themethod of claim 1, wherein the silicon substrate includes asemiconductor-on-insulator (SOI) structure including anelectrically-insulating layer, the method further including wet etchingthe buried oxide of the SOI structure.
 8. The method of claim 1, furtherincluding forming the fluidic channel by bonding a fluidic structure tothe silicon substrate.
 9. The method of claim 8, wherein the siliconsubstrate include a semiconductor-on-insulator (SOI) structure includingan electrically-insulating layer and forming the fluidic channelincludes placing and bonding the fluidic structure such that the fluidicchannel is defined by a portion of an upper-facing surface of theelectrically-insulating layer of the SOI structure with the definedlocation for the fluidic channel and by a bottom-facing surface of thefluidic structure.
 10. The method of claim 1, further including formingat least two electrodes, on the silicon substrate, configured andarranged to pass current for facilitating electrophoretic flow orelectroosmotic flow of the fluid through the fluidic channel.
 11. Themethod of claim 10, further including implanting boron at the end of thesemiconductor regions for ohmic contact to the at least two electrodes.12. The method of claim 10, further including adding an amplificationcircuit supported by the silicon substrate and configured and arrangedto facilitate sensing the target material near or in the fluidicchannel, wherein in response to the potential resulting from the targetmaterial in the fluidic channel, the at least one semiconductor deviceis configured and arranged to sense the presence of the target materialin the fluidic channel in response to the flow of electrons in aconducting mode from one end electrode to another end electrode.
 13. Amethod of manufacturing an apparatus for sensing a target material, themethod comprising: patterning at least one semiconductor device onto asilicon substrate, wherein the at least one semiconductor deviceincludes at least three electrically-contiguous semiconductor regionsand end electrodes; doping the at least three electrically-contiguoussemiconductor regions to exhibit a common polarity, wherein the at leastthree electrically-contiguous semiconductor regions includes asandwiched region sandwiched between two of the at least threeelectrically-contiguous semiconductor regions, wherein the two of the atleast three electrically-contiguous semiconductor regions are doped at ahigher concentration than the sandwiched region and are electricallyconnected to the end electrodes; defining a location, on the siliconsubstrate for a fluidic channel under the at least one semiconductordevice, wherein the semiconductor regions are configured and arrangedadjacent to the fluidic channel with a surface directed toward thefluidic channel for coupling to the target material in a fluid in thefluidic channel, and for responding to a change in potentialattributable to the target material; and forming the fluidic channel bybonding a fluidic structure of the fluidic channel to the siliconsubstrate such that the fluidic channel is defined by a portion of anupper-facing surface of the silicon substrate and by a bottom-facingsurface of the fluidic structure.
 14. The method of claim 13, whereinforming the fluidic channel further includes: pouring a material into amold to form the fluidic structure; and once solidified, separating thefluidic structure from the mold.
 15. The method of claim 14, furtherincluding exposing the fluidic structure to oxygen plasma.
 16. Themethod of claim 13, further including placing the apparatus in a heatingdevice for a period of time to adhere the fluidic structure to thesilicon substrate.
 17. The method of claim 13, further including:forming at least two electrodes, on the silicon substrate, configuredand arranged to pass current for facilitating electrophoretic flow orelectroosmotic flow of the fluid through the fluidic channel; andforming an amplification circuit supported by the silicon substrate andconfigured and arranged to facilitate sensing the target material nearor in the fluidic channel.
 18. The method of claim 17, wherein inresponse to the potential resulting from the target material in thefluidic channel, the at least one semiconductor device is configured andarranged to sense the presence of the target material in the fluidicchannel in response to the flow of electrons in a conducting mode fromone end electrode of the end electrodes to another end electrode of theend electrodes.
 19. The method of claim 13, wherein the target materialincludes biological molecules.
 20. The method of claim 13, wherein thetarget material includes chemical molecules.